The accurate deposition and patterning of biological probes on a solid support is of critical importance to numerous bioassays. For example, protein and DNA microarrays, which offer the interesting possibility to study concurrently the interaction between a target sample and a large number of different biological probes, have become key components of drug discovery, clinical diagnostics, and gene sequencing. However, microarrays still largely depend on detection techniques such as fluorescence labeling or surface plasmon resonance which are difficult to apply in point-of-care applications. Recently, new detection techniques involving integrated sensors with microfabricated biosensing elements have emerged, including: nanowires, field-effect-transistors, optical sensing waveguides, and electrochemical sensors. This new generation of microfabricated biosensor arrays creates a pressing need for the development of techniques that allow the high-quality immobilization of various biological probes with high positional accuracy on the micron-size sensing elements of the chips.
Numerous techniques have been developed for the immobilization of DNA, proteins, cells or other biological probes on a solid surface, including: pin printing, inkjet printing, microstamps, and microfluidics. Pin printing, in which solid metal pins are pressed on a surface to transfer minute amount of liquid, is still a widely used technique due to its relative simplicity and the possibility to pattern arrays with thousands of spots. However, accurate positioning and registration of the spots are difficult to control and require costly and sophisticated tools. Also, rapid and uncontrolled drying of the liquid deposited can lead to non-uniform spots and denaturing conditions, especially when the dimensions of the spots are decreased below about 80 μm.
Microfluidics provides a simple path to better control the immobilization conditions as well as the dimension, positioning, and uniformity of the deposition zone. Microfluidic immobilization devices generally consist of a network of channels patterned in polydimethylsiloxane (PDMS), a thermoset elastomer that can create a reversible conformal sealing to most solid supports. The biological probes are then flown in the device, incubated, and washed, thus giving rise to immobilized patterns matching the geometry of the channels. With this technique, the dimension of the spots is precisely set by the geometry of the channels (spot width of less than 1 μm has been demonstrated) and better control of the immobilization conditions is achieved, which is of critical importance for sensitive biomolecules such as proteins. On the other hand, simple microfluidic devices having a 2D network of channels are inherently limited to pattern continuous features and cannot be used to form an array of isolated spots, as it would be required, for example, to immobilize biological probes only on the sensing elements of a microfabricated biosensor array. Also, because the biological probes are in contact with the activated substrate over the entire length of the 2D microfluidic device, a rapid depletion of the immobilization solution is typically observed during the transport in the channels.
To overcome these limitations, various more complex designs based on a 3D geometry have been proposed. In these designs, the channels are typically embedded inside the microfluidic device and the liquid is brought in contact with the substrate only on the desired locations using open-through holes (e.g. “vias”) (Chiu 2000; Griscom 2001; Juncker 2002; Juncker 2005; Kloter 2004; Wang 2006), channels oriented perpendicularly to the substrate (Chang-Yen 2006; Eddings 2008; Natarajan 2008a; Natarajan 2008b; Eddings 2009), or flow confinement effects (Hofmann 2002; Juncker 2005; Eddings 2009).
3D microfluidic immobilization devices have first been demonstrated by Chui et al. who reported the patterning of up to three types of proteins or cells on isolated regions by using a thin PDMS membrane with open-through holes to make connections between two layers of channels (Chiu 2000). Junker et al. also reported 3D microfluidic devices made by etching open-through holes in silicon wafers with deep reactive ion etching (Juncker 2002; Juncker 2005; Kloter 2004). Capillary phenomena were then used as the driving force to pattern of up to 11 independent 50 μm size protein spots on a PDMS substrate. Recently, Gale's and Myszaka's groups designed a multi-layer 3D patterning system based on channels oriented perpendicularly to the substrate allowing up to 48 independent biological probes to be immobilized on isolated spots of about 400 μm size (Chang-Yen 2006; Eddings 2008; Natarajan 2008a; Natarajan 2008b; Eddings 2009).
Sudarsan et al. describes the fabrication of microfluidic devices made of a TPE consisting of a home-developed mixture of SEBS and mineral oil prepared by heating the constituents in vacuum overnight (Sudarsan 2004a, Sudarsan 2004b, Sudarsan 2005). Although a device made from few layers of this TPE was briefly described, no method was unveiled to fabricate devices with a dense array of vias with TPE or to create microscopic open-through holes in such material. In the only example presented, the fluidic connection between the layers was made by simply punching a macroscopic hole manually. This method cannot be used to create complex devices with 3D network of channels.
Stoyanov et al. used thermoplastic polyurethane foils TPU (a specific type of TPE) to fabricate solvent resistant microfluidic devices for the use with a surface acoustic wave sensor chip (Stoyanov 2005; Stoyanov 2006). The foils were patterned by hot embossing and open-through holes were demonstrated. However, no 3D microfluidic devices were demonstrated or discussed. Only two rather large (greater than 300 μm) open-through holes were patterned on the devices and they were used only as an inlet and outlet (not as 3D interconnects or vias). Also the TPU grade used for the experiment (Walopur™ 2201 AU) has shore hardness higher than 85 A, which is too high to provide reversible and conformal sealing on a surface or to allow the demolding of undercut profiles. Also, a high pressure (50-120 bar) was needed to correctly pattern the TPU foils, which prevented the use of low cost photoresist molds and required metallic molds.
Despite these recent developments, many challenges must still be solved before microfluidics can be accepted as a universal biological patterning tool. Existing processes are typically very challenging, labor-intensive, and/or inherently serial. Also, the properties of the materials used in such 3D devices are typically far from ideal. As a consequence, the compatibility of the devices is often limited to only the most standard solvents and complicated and costly steps are required for the patterning and bonding of the multiple layers from which the devices are built. These drawbacks have relegated 3D microfluidics to relatively simple academic prototyping and have largely dissuaded researchers and industries from further research on these methods. As a consequence, despite more than ten years of research since the concept of 3D microfluidics was first demonstrated, microfluidic devices are still today almost exclusively based on network of channels patterned on a single 2D plane.
The most critical issue arguably arises from the intrinsic need to use the microfluidic patterning devices only once to avoid cross-contamination issues. Under such circumstances, the development of high-throughput mass-production processes to achieve low-cost per device is of critical importance. Unfortunately, almost all previous designs of microfluidic patterning devices have relied on PDMS, which is not very amendable to low-cost mass-production. Other drawbacks of PDMS are discussed below. It is also noteworthy that many of the 3D microfluidic designs proposed to date require the precise and difficult alignment of many elastomeric layers over large areas and the use of costly and lengthy post-processing procedures to punch the numerous access holes and cut the devices in final shape.
Approaches to fabricating 3D microfluidic devices can be divided into three categories: (i) layered PDMS microfluidics; (ii) layered microfluidic devices made from hard materials; and, (iii) 3D molding.
Layered PDMS Microfluidics
The most common method of achieving 3D devices involves the fabrication and stacking of several thin open-through layers of PDMS (polydimethylsiloxane), a soft thermoset elastomer. Each individual layer is fabricated either by spin casting uncured PDMS prepolymer on a mold (so that the highest features of the mold breach through the PDMS layer) or by clamping a drop of PDMS prepolymer between a mold and a top plate. The thin PDMS layers are then cured, peeled off from the mold, oxidized in O2 plasma, aligned, and bonded into a 3D microfluidic device.
PDMS is the standard and most widely used material for both 2D and 3D microfluidics. Although it has some very attractive properties such as high transparency, low hardness, elasticity, and relatively low cost, it also shows some serious drawbacks, which have precluded industry adoption of PDMS for mass production. Firstly, as PDMS is a thermoset, it requires lengthy curing and degassing steps, which makes its use very unpractical for mass production. This problem becomes critical for the fabrication of layered 3D devices, as many layers need to be degassed and cured independently for the fabrication of a single device. The thermosetting properties of PDMS also prevent the use of simple techniques such as thermal bonding to assemble the final devices. Indeed, the bonding of PDMS layers typically involves a plasma oxidation step that must be rapidly (less than 1 min) followed by the alignment and bonding of the layers. Another problem is the intrinsic porosity of PDMS and its relatively high gas permeability. As a result, water tends to evaporate quickly through PDMS, which limits the maximum length of an assay and can be critical for applications where osmolality must be carefully monitored (e.g., cellular studies, etc.). This problem is also strongly exacerbated in 3D layered devices due to the use of thin layers of PDMS (typically about 100 μm).
The fabrication of layered 3D devices from PDMS also typically requires manual peeling of the membrane from the molds. This process is not only inherently serial but is also very problematic due to the rather low mechanical strength of PDMS. The PDMS membranes can thus break or deform very significantly during their manipulation, which makes alignment difficult or even impossible. Finally, PDMS is not compatible with a large number of solvents and is thus relegated mostly to water-based chemistry. For example, PDMS will absorb not only hydrocarbon solvents but also some analytes with a slight lipidic character.
Thus, although PDMS is very relevant for prototyping and academic demonstration of concepts, its use for the mass production of complex 3D layered devices is far from ideal.
Layered Microfluidic Devices Made from Hard Materials
A similar layered approach has also been used to fabricate 3D microfluidic devices from the following hard materials: silicon, glass, ceramic, metal, hard thermoplastics, biodegradable polymers, photo curable polymers, photoresists, and paper. Although the fabrication techniques vary greatly depending on the material of interest, they all involve the production of open-through layers and their bonding into a 3D device. Depending on the material, the open-through holes have been obtained by techniques such as drilling, etching, punching, photopatterning, hot embossing, and laser cutting. The layers then have to be bonded into functional 3D devices by using techniques such as thermal bonding, chemical bonding, photoresist curing, or double-sided adhesive tape.
The use of hard materials for layered 3D microfluidics can alleviate some of the problems encountered with PDMS. They however have their own limitations. The most important drawback comes from the rigid nature of these materials. Contrary to elastomeric soft materials, hard materials do not allow reversible and conformal sealing on an arbitrary surface and do not offer the possibility of creating easily implementable valving schemes. Some of materials involved in the fabrication of 3D devices are also not transparent (silicon, ceramic, wax, paper, etc.).
Fabrication of multi-layers 3D devices with hard materials is also typically more problematic than with PDMS. The patterning of inorganic materials such as glass, silicon, metal, and ceramics cannot be performed with low cost rapid prototyping tools. The production costs with these materials are thus generally too high to produce single-use complex 3D devices at reasonable price. The patterning of hard thermoplastics is generally much easier than for hard inorganic materials, but it also presents some issues. It is indeed very challenging to create microscopic open-through holes reliably in hard thermoplastics. For example, with hot embossing, both high pressures and temperatures are required to correctly transfer the pattern of the mold and to punch through a plastic sheet. Under these conditions, it is not possible to use low cost photoresist-based molds (as typically used in PDMS molding), and costly metallic molds must be prepared. The rigid nature of hard thermoplastics also makes demolding difficult or even impossible when high aspect ratio features are required. Various strategies must also be implemented to avoid the presence of a thin residual layer at the top of each open-through hole. The typical scheme requires the use and alignment of a receptor mold with holes corresponding to the protruding features of the embossing mold. It is also possible to use a polymeric sacrificial layer so that the mold features protrude in this second layer and leaves open-through holes in the thermoplastic part.
Finally, it must be stressed that bonding is a difficult problem for most hard materials. It is generally achieved by pressing together the various layers under a specific force and temperature. However, due to the high rigidity of these materials, microscopic defects, surfaces irregularities, or non ideal bonding conditions can easily result in partially bonded section and leaks. As the probability of defect increases with the number of bonded layers, it can be very challenging to produce complex 3D devices in a reliable manner from hard materials. Consequently, bonding techniques are still a very active research area in 3D microfluidics. For example, approaches using double sided tape, or partially cured photoresists have been recently proposed to improve the bonding reliability. Nevertheless, due to their rigid nature, hard materials are not compatible with applications requiring reversible and conformal sealing on arbitrary surfaces.
3D Molding and Direct 3D Fabrication
The third and last approach involves the fabrication of a 3D sacrificial mold containing directly the desired final geometry for the network of channels. The microfluidic device is then fabricated from this 3D mold by using techniques such as metal electroforming or casting of a prepolymer or an epoxy. The final microfluidic device is then released by melting or dissolving the mold. It is to be noted that the mold has to be sacrificed and cannot be reused as soon as the design contains suspended features. The molds are generally fabricated by solid object printing of low fusion temperature materials such as wax. Alternatively, the microfluidic devices can also be fabricated directly by 3D fabrication techniques similar to that used for the fabrication of the 3D sacrificial molds (e.g. stereolithography).
The main advantage of 3D molding and direct 3D fabrication is to eliminate the alignment and bonding steps required in layered fabrication. However, the lengthy and costly process of fabricating either a complex sacrificial 3D mold for each device or each 3D microfluidic device in a serial manner limits this technique to device design and early prototyping. The printed 3D molds also typically have high roughness and a relatively low resolution (about 100 microns). Finally, some types of features, such as long suspended channels, are difficult to create with this approach.
Although proof of concept for microfluidic immobilization of biological probes was obtained more than a decade ago, none of the microfluidic devices proposed to date can clearly combine all the characteristics necessary for widespread adoption of the technology. There remains a need for new methods to build complex 3D microfluidic devices using simple techniques and materials that have appropriate properties for the targeted applications.